In past nuclear imaging devices, gamma radiation detectors employ scintillators that convert incident gamma radiation into light, which is then detected by photomultiplier tubes (PMTs). A scintillator is a material which exhibits scintillation - the property of luminescence when excited by ionizing radiation. Due to several drawbacks of photomultiplier tubes, there is interest in replacing them with solid state light sensors, such as avalanche photodiodes driven in Geiger mode, called e.g., silicon photomultipliers (SiPMs). Typical SiPMs have better timing and energy resolution than typical PMTs. A timing resolution significantly better than one nanosecond is becoming more valuable as time-of-flight PET (TOF-PET) scanners are becoming more prevalent, however, there have been serious impediments to adopting this new technology.
A pixellated detector is composed of a high resistivity semiconductor chip containing pixellated photodiodes with their respective readout electronics. This semiconductor chip is also called diode detection layer. For high sensitivity, the ratio of the diode area to the full area per pixel, called fill factor, should be high, typically above 50%. X-rays are absorbed in scintillator crystals that are on top of and optically coupled to the diode detection layer. The optical photons generated in the scintillator crystals are detected by the diodes of the respective individual pixels in the diode detection layer and converted into electrical signals. The signal of each pixel diode is read out by a specific read-out electronics channel on the semiconductor chip.
A gamma camera, also called a scintillation camera or Anger camera, is a device used to image gamma radiation emitting radioisotopes, a technique known as scintigraphy. The applications of scintigraphy include early drug development and nuclear medical imaging to view and analyse images of the human body or the distribution of medically injected, inhaled, or ingested radionuclides emitting gamma rays. Current SPECT detectors and early PET detectors have been built based on such an Anger camera with a continuous NaI:Tl crystal. Modern PET detectors use either a block detector, or arrays of individual scintillator crystals which are optically separated from each other with a reflective material. A suitable scintillator for TOF-PET is LYSO (Lu1.8Y0.2SiO5:Ce), a suitable reflection layer can be obtained by wrapping the crystals in a layer of Teflon. Such an array is optically coupled to an array of PMTs, using an intermediate ‘light guide’ layer to spread the light generated by a gamma quantum in an individual scintillator crystal over the PMT array so that it is possible to use Anger logic.
Newer generations of PET detectors use much smaller detector pixels implemented as silicon photomultipliers (SiPM). The concept has generally been based on one-to-one coupling of scintillator crystal and SiPM. The idea is to measure the light generated within one scintillator pixels with only one SiPM detector, in order to maximize the signal on this detector, and to minimize the data readout rate and the influence of detector dark counts on the signal. Dark counts are an inherent property of SiPM technology. Accounting for even low optical crosstalk to neighboring pixels or for Compton scatter, at least nine detector pixels would have to be read out, the ‘direct’ detector pixel plus its eight neighbors. This larger readout area would require a nine-fold readout rate, and would mean a significantly larger contribution of dark counts to the signal. A promising concept for reflectors in scintillator arrays is the use of reflective sheets, e.g. Vikuiti Enhanced Specular Reflectors (Vikuiti ESR). These dielectric mirrors provide high reflectivity, very low optical crosstalk and no optical absorption, and they enable a high fill factor due to their thickness of only 65 μm. A part of the scintillation light is however channeled in the gap between crystal and reflector, yielding a higher light output right along the edges of the pixel. This increased light output from the interface just between crystal and reflector can have the result that, in spite of the small interface area, the scintillation light from this area contributes about 10-20% of the total signal. In a typical one-to-one coupling, however, the sensitive area of each detector pixel is centered below one scintillator crystal, while the non-sensitive areas of the chip (readout electronics etc.) are placed below the ‘gaps’ between the crystals which means that such a detector exactly misses most of the area of the crystal where the highest light output occurs. In addition, light from these ‘gaps’ can instead reach neighboring detector pixels, thereby increasing the undesirable light crosstalk.
FIG. 1 shows a schematic top view of a conventional pixellated detector device with an arrangement of scintillator crystals 50 with reflector cover 30 in one-to-one correspondence to detector pixels 10 with active photosensitive area.
FIGS. 2A and 2B show cross-sections of the conventional pixellated detector devices as shown in FIG. 1 without (FIG. 2A) and with a common glass substrate 60 (FIG. 2B). The arrows show light emitted from edge regions of scintillator pixels. Crosstalk would be high and part of the light would be lost with the common glass substrate 60 between crystals 50 and detector pixels 10 (FIG. 2B), even if the reflector sheets should be 100% reflective. Structuring the glass plate might be a solution but is risky and expensive.
Consequently, in the above conventional technical approach based on a one-to-one coupling of scintillator pixels and detector pixels (e.g. SiPM pixels) signal is lost, and it is extremely difficult to avoid optical crosstalk between pixels.